Atherosclerosis is an important cause of death in many parts of the world. Therefore, many techniques have been developed to obtain information about the plaque that develops in blood vessels. Image techniques such as angiography, magnetic resonance imaging, intravascular ultrasound, and optical coherence tomography provide information regarding the location of a plaque or blood vessel obstruction and about the morphology or internal structure of the plaque. However, they do not enable detailed in vivo analysis of the molecular composition of the plaque. Knowledge of the molecular composition of a plaque is important e.g. for determining the risk of acute cardiac events. So-called stable plaque and vulnerable plaque are distinguished, where it is thought that the vulnerable plaque can give rise to such acute, often fatal events. Such an event is triggered by rupturing of the thin fibrous cap of the plaque, bringing the contents of the lipid pool of the plaque into contact with the blood stream, leading to thrombogenesis and occlusion of the artery.
Fluorescence based methods have been shown to be able to distinguish between normal artery wall and atherosclerotic plaque in vitro. However fluorescence spectra arc easily disturbed by light absorbing molecules in the tissue and in blood, limiting its applicability.
Of all methods to obtain information about atherosclerotic plaque composition and which can in principle be applied in vivo, intravascular Raman spectroscopy provides the most detailed information. In Raman spectroscopy, the Stokes-shift between light that is incident on a sample that is investigated and the light that is in elastically scattered by the sample is expressed in relative wavenumbers (Δcm−1=(1/λin−1/λscattered)10−2 with λ (wavelength) in meter). The wavenumber region between about 400 and 2000 cm−1 of the Raman spectrum (the so-called fingerprint region) is used to obtain this information. This region of the spectrum contains many bands that can be discerned and which individually and/or in combinations can be used to obtain information about the molecular composition of the tissue.
Studies in the field of atherosclerosis are only related to the fingerprint region, since this spectral region is very informative for analysis or diagnosis. Examples of such studies an e.g. found in the papers of H. P. Buschman, E. T. Marple, M. L. Wach, B. Bennett, T. C. Schut, H. A. Bruining, A. V. Bruschke, A. van der Laarse. and G. J. Puppels, Anal. Chem. 72 (2000), 3771-3775, which discusses the in vivo determination of the molecular composition of artery wall by intravascular Raman spectroscopy, using a multifiber probe and measuring in the 400-1800 cm−1 region; R. H. Clarke, E. B. Hanlon, J. M. Isner, H. Brody, Appl. Optics 26 (1917), 3175-3177, which discusses laser Raman spectroscopy of calcified atherosclerotic lesions in cardiovascular tissue, also in the fingerprint region; and J. F. Brennan T. J. Romer, R. S. Lees, A. M. Tercyak, J. R. Kramer, M. S. Feld, Circulation 96 (1997), 99-105, which deals with the determination of human coronary artery composition by Raman spectroscopy in the fingerprint region.
In vivo application of Raman spectroscopy in most cases require the use of a flexible light guiding device of small diameter. This can for instance be introduced in the lumen of an artery. It must be able to reach and interrogate locations with atherosclerotic lesions. It can also be used in the working channel of an endoscope or inside a biopsy needle or biopsy forceps. The fiber optic probe (comprising one of more optical fibers) must guide light to the tissue under investigation, collect light that is scattered by the tissue and transport this collected light away from the tissue towards a spectrum analysis device.
Unfortunately, in the 400-2000 cm−1 spectral region, the materials of the optical fiber itself generate Raman signal, resulting in a strong signal background. Moreover, bending of the fiber leads to variations in the amount of signal obtained from the core, cladding and coating materials, further complicating signal detection and signal analysis. This deteriorates the signal-to-noise with which the tissue Raman signal can be detected, and also otherwise complicates signal analysis, and therefore negatively affects clinical utility. It is therefore necessary to use optical filters at or near the distal end of the fiber optic probe which is in contact or in close proximity to the tissue, in order to suppress background signal contributions to the detected tissue Raman signal. However, this in turn necessitates the use of separate optical fibers for guiding laser light to the tissue and for collecting and guiding scattered light away from the tissue. It furthermore often necessitates the use of beam steering arrangements or a lens or lenses at the distal end of the fiber optic probe in order to obtain the desired overlap between the volume of tissue illuminated by the laser light and the volume of tissue from which Raman signal can be collected. Fiber optic probes for Raman spectroscopy are therefore complex. It is difficult to miniaturise fiber optic probes for Raman spectroscopy and to keep them flexible, which is necessary for instance for intravascular use and for use in the auxiliary channel of an endoscope. The complexity is also an obstacle to the production of such probes at a price that they can be used as disposables in hospitals. Moreover, signal intensity of tissue in the 400-2000 cm−1 is low, necessitating relatively long signal integration times, which may be impractical for clinical use. All above mentioned problems and disadvantages hinder the actual implementation of Raman spectroscopy for clinical diagnostic purposes in general, and for intravascular use in particular.
Light is guided through the optical fiber in so-called bound modes. In these bound modes the electromagnetic field is located primarily in the core of the optical fiber, with a small part extending into the cladding. Lower order modes are more confined to the core than higher order modes.
The intensity of light that is guided by an optical fiber is attenuated. This is caused by absorption, by light scattering (Rayleigh scattering, scattering/reflection at larger inhomogeneities or at sites at which the fiber material is damaged), and by micro and macro-bending losses.
Laser light, that is lost by scattering events leaves the core of the fiber and passes through the coating (and buffer) layers. Coating and buffer layers are usually made of silicon or plastic or polymer material.
U.S. Pat. No. 5,293,872 teaches the use of near-infrared (NIR) laser light excited Raman spectroscopy for distinguishing between normal artery tissue, calcified atherosclerotic plaque and fibrous atherosclerotic plaque. For in vivo measurements in the 700-1900 cm−1 region, the use of a bundle of optical fibers is discussed. This will lead to the same disadvantages as discussed above, e.g. with respect to noise.
U.S. Pat. No. 5,194,913 recognises the problem of multiple fiber optics, but also notes that the use of a single fiber is prohibited by the fact that background Raman signal generated in the fiber optics is intense for all but the shortest fibers. It discloses a fiber optic apparatus using two opposite fibers and using optical filters to reduce background Raman emission from the fiber optics. This document is related to the problem of signals in fibers in general, and it is clear that the solution provided by U.S. Pat. No. 5,194,913, i.e. an axial configuration, cannot easily be used for measurements in vivo.
A paper of J. F. Aust, K. S. Booksh and M. L. Myrick, Applied Spectroscopy 50 (1996), 382-386 discusses cases in which the signal obtained from the sample is relatively strong (polymer) or in which special measures were taken, such as increasing the measurement volume from which sample-Raman signal is obtained, to increase signal intensity from polymers to levels that are very much higher than would be obtained from a biological tissue. This paper does not discuss the applicability of the method to tissue, but teaches that for a good signal, a special Teflon tube of up to 4 cm has to be used on the tip of the optical probe, filled with the polymer, in order to get a good signal. Such a method is usually not applicable to tissue, especially not in the case of in vivo measurements.
Next to atherosclerosis, cancer is also an import health issue. The same problems as encountered above apply for determining tumor cells by Raman spectroscopy via fiber optics. U.S. Pat. No. 5,261,410 teaches to use a bundle of fibres and to measure in the fingerprint region. Such use also leads to a signal to noise ratio which is not satisfying.
From the above it is clear that there is a need for an cut for an instrument for measuring a Raman signal of a tissue, that does not have above mentioned problems.